Real-time volumetric bi-plane ultrasound imaging and quantification

ABSTRACT

Quantified measures of a volumetric object in the body can be made ultrasonically by acquiring concurrent biplane images of two different image planes of the object. Corresponding borders of the volumetric object are traced using automatic border detection. The border tracings are used in their planar spatial relationship to compute a graphical model of the volumetric object. The volume of the graphical model may be computed by the rule of disks, and a graphical or numerical display of the changing volume with time displayed. A user interface comprises both real time biplane images, the real time graphical model, and the quantified measures.

CROSS REFERENCE TO RELATED APPLICATION

This application claims the benefit of U.S. provisional application Ser.No. 60/507,263 filed Sep. 29, 2003, which is incorporated herein.

This invention relates to ultrasonic diagnostic imaging and, moreparticularly, to ultrasonic imaging systems capable of estimating thevolume of vessels and organs such as the heart.

Echocardiographic ultrasonic imaging systems are used to assess theperformance of the heart. Cardiac performance can be assessedqualitatively with these systems, such as by observing the blood flowthrough vessels and valves and the operation of heart valves.Quantitative measures of cardiac performance can also be obtained withsuch systems. For instance, the velocity of blood flow and the sizes oforgans and cavities such as a heart chamber can be measured. Thesemeasures can produce quantified values of cardiac performance such asejection fraction and cardiac output. A method and apparatus formeasuring the volume of a heart chamber are described in U.S. Pat. No.5,322,067 (Prater et al.), for example. In the method described in thispatent, the clinician acquires a sequence of ultrasound images of acavity to be measured, such as the left ventricle of the heart. Theclinician freezes one of the images on the display screen and traces afixed region of interest (ROI) around the cavity of the heart chamber.The defined ROI should be large enough to encompass the heart chamberwhen the heart is fully expanded. The ultrasound system then processesthe pixels in the ROI in each image in the sequence to determine thosepixel that are blood pixels in the left ventricle. Each left ventricleis then segmented into strips and the area of the strips is calculated.Each strip is then conceptually rotated about its center to define adisk and the volume of each disk is calculated. By summing the volumesof the disks in each image the volume of the heart chamber is determinedat each point in the heart cycle for which an image was acquired. Thecalculated volumes can be displayed numerically as a function of time,or a waveform representative of left ventricle volume as a function oftime can be produced, thereby showing the clinician the changes in leftventricular volume over the heart cycle.

The method of the Prater et al. patent requires manual input from theclinician who must define the ROI by a manual tracing. The method canonly be performed on a stored image loop due to this need for manualinput. It would be desirable for such a technique to be performed by theultrasound system automatically and to be performed in true real time asthe images are acquired. Furthermore, the method of disks (Simpson'srule) volume estimation assumes that each disk is uniformly circular,which may not be the case. It would be desirable to estimate cavityvolumes that are more closely related to the true shape of the anatomyrather than having to rely on an assumption of geometric uniformity ofthe anatomy, thus producing more accurate volume measures.

In accordance with the principles of the present invention, the volumeof a body cavity or organ is measured by ultrasonic imaging.Cross-sectional images in different planes of the body cavity areacquired at substantially the same time, thus presenting views of theshape of the cavity at a point in time from different perspectives. Asurface of the cavity or organ in each image is outlined by automated orsemi-automated border tracing. Segments of the cavity are defined byproducing a geometric model of the cavity or organ from the tracings.The segment volumes are accumulated to produce an accurate measure ofthe volume of the cavity. The inventive method can be performed in realtime and produces a more accurate estimate of the cavity volume. Theresulting measure can be displayed numerically or as a physiologic curveof the changing volume with time.

In the drawings:

FIG. 1 is a four chamber ultrasound image of the heart.

FIG. 2 illustrates an ultrasound display of both end diastole and endsystole cardiac images.

FIGS. 3 a and 3 b illustrate the step of locating the medial mitralannulus (MMA) and the lateral mitral annulus (LMA) in an ultrasoundimage of the left ventricle (LV).

FIG. 4 illustrates the step of locating the apex of the LV.

FIGS. 5 a-5 c illustrate standard border shapes for the LV.

FIGS. 6 a-6 b illustrate geometric templates used to locate the MMA andLMA.

FIGS. 7 a-7 c illustrate a technique for fitting a standard border shapeto the endocardial boundary of the LV.

FIG. 8 illustrates an end diastole and end systole display withendocardial borders drawn automatically in accordance with theprinciples of the present invention.

FIG. 9 illustrates the rubber-banding technique for adjusting anautomatically drawn border.

FIG. 10 is a photograph of an actual ultrasound system display whenoperating in the biplane mode in accordance with the principles of thepresent invention.

FIG. 11 illustrates in block diagram form an embodiment of an ultrasonicdiagnostic imaging system constructed in accordance with the principlesof the present invention.

FIG. 12 illustrates an ultrasound display screen produced in accordancewith the principles of the present invention.

FIGS. 13 a and 13 b illustrate the formation of cavity segments fromorthogonal views of the left ventricle.

Referring first to FIG. 1, an ultrasound system display is shown duringthe acquisition of cardiac images. The ultrasound image 10 is afour-chamber view of the heart which is acquired by a phased arraytransducer probe to produce the illustrated sector-shaped image. Theimage shown is one of a sequence of real-time images acquired byplacement of the probe for an apical 4-chamber view of the heart, inwhich the probe is oriented to view the heart from the proximity of itsapex 11. The largest chamber in the image, in the central and upperright portion of the image, is the left ventricle (LV). As the real-timeultrasound image sequence is acquired a scrolling ECG trace 12 of theheart cycle is simultaneously acquired and displayed at the bottom ofthe display, with a triangular marker 14 denoting the point or phase ofthe cardiac cycle at which the currently-displayed image was acquired. Atypical duration of the heart cycle when the body is at rest is aboutone second, during which time approximately 30-90 image frames of theheart can be acquired and displayed in rapid succession. As theclinician views the display of FIG. 1, the heart is seen beating in realtime in the ultrasound display as the ECG waveform 12 scrolls beneaththe ultrasound images 10, with the instantaneously displayed heart phaseindicated by the marker 14.

In one mode of acquisition, the clinician observes the beating heart inreal time while manipulating the transducer probe so that the LV isbeing viewed distinctly in maximal cross-section. When the four chamberview is being acquired continuously and clearly, the clinician depressesthe “freeze” button to retain the images of the current heart cycle inthe image frame or Cineloop® memory of the ultrasound system. TheCineloop memory will retain all of the images in the memory at the timethe freeze button is depressed which, depending upon the size of thememory, may include the loop being viewed at the time the button wasdepressed as well as images of a previous or subsequent loop. A typicalCineloop memory may hold 400 image frames, or images from about eight toten heart cycles. The clinician can then scan through the stored imageswith a trackball, arrow key, or similar control to select the loop withthe images best suited for analysis. When the clinician settles on aparticular loop, the “ABD” protocol is actuated to start the borderdrawing process.

When the ABD protocol is actuated the display changes to a dual displayof the end diastole image 16 and the end systole image 18 displayedside-by-side as shown in FIG. 2. The ultrasound system identifies all ofthe images comprising the selected loop by the duration of the ECGwaveform associated with the selected loop. The ultrasound system alsorecognizes the end diastole and end systole points of the cardiac cyclein relation to the R-wave of the ECG waveform 12 and thus uses the ECGwaveform R-wave to identify and display the ultrasound images at thesetwo phases of the heart cycle. The dual display of FIG. 2 shows the ECGwaveform 12 for the selected heart cycle beneath each ultrasound image,with the marker 14 indicating the end diastole and end systole phases atwhich the two displayed images were acquired.

Since the Cineloop memory retains all of the images of the cardiaccycle, the user has the option to review all of the images in the loop,including those preceding and succeeding those shown in the dualdisplay. For instance, the clinician can “click” on either of the imagesto select it, then can manipulate the trackball or other control tosequentially review the images which precede or succeed the one selectedby the ultrasound system. Thus, the clinician can select an earlier orlater end diastole or end systole image from those selected by theultrasound system. When the clinician is satisfied with the displayedimages 16 and 18, the ABD processor is actuated to automaticallydelineate the LV borders on the two displayed images as well as theintervening undisplayed images between end diastole and end systole.

In this example the ABD processor begins by drawing the endocardialborder of the LV in the end systole image 18. The first step in drawingthe border of the LV is to locate three key landmarks in the image, themedial mitral annulus (MMA), the lateral mitral annulus (LMA), and theendocardial apex. This process begins by defining a search area for theMMA as shown in FIG. 3 a, in which the ultrasound image grayscale isreversed from white to black for ease of illustration. Since the ABDprocessor is preconditioned in this example to analyze four-chamberviews of the heart with the transducer 20 viewing the heart from itsapex, the processor expects the brightest vertical nearfield structurein the center of the image to be the septum which separates the left andright ventricles. This means that the column of pixels in the image withthe greatest total brightness value should define the septum. With thesecues the ABD processor locates the septum 22, and then defines an areain which the MMA should be identified. This area is defined fromempirical knowledge of the approximate depth of the mitral valve fromthe transducer in an apical view of the heart. A search area such asthat enclosed by the box 24 in FIG. 3 a is defined in this manner.

In this embodiment a filter template defining the anticipated shape ofthe MMA is then cross correlated to the pixels in the MMA search area.While this template may be created from expert knowledge of theappearance of the MMA in other four-chamber images as used by Wilson etal. in their paper “Automated analysis of echocardiographic apical4-chamber images,” Proc. of SPIE, August, 2000, a geometric cornertemplate may be used as follows. While a right-angle corner template maybe employed, in a constructed embodiment an octagon corner template 28(the lower left corner of an octagon) is used as the search template forthe MMA, as shown at the right side of FIG. 6 a. In practice, theoctagon template is represented by the binary matrix shown at the leftside of FIG. 6 a. The ABD processor performs template matching by crosscorrelating different sizes of this template with the pixel data indifferent translations and rotations until a maximum correlationcoefficient above a predetermined threshold is found. To speed up thecorrelation process, the template matching may initially be performed ona reduced resolution form of the image, which highlights majorstructures and may be produced by decimating the original imageresolution. When an initial match of the template is found, theresolution may be progressively restored to its original quality and thelocation of the MMA progressively refined by template matching at eachresolution level.

Once the MMA has been located a similar search is made for the locationof the LMA, as shown in FIG. 3 b. The small box 26 marks the locationestablished for the MMA in the image 18, and a search area to the rightof the MMA is defined as indicated by the box 34. A right cornergeometric template, preferably a right octagon corner template 38 asshown in FIG. 6 b, is matched by cross-correlation to the pixel valuesin the search area of box 34. Again, the image resolution may bedecimated to speed the computational process and different templatesizes may be used. The maximal correlation coefficient exceeding apredetermined threshold defines the location of the LMA.

With the MMA 26 and the LMA 36 found, the next step in the process is todetermine the position of the endocardial apex, which may be determinedas shown in FIG. 4. The pixel values of the upper half of the septum 22are analyzed to identify the nominal angle of the upper half of theseptum, as indicated by the broken line 43. The pixel values of thelateral wall 42 of the LV are analyzed to identify the nominal angle ofthe upper half of the lateral wall 42, as shown by the broken line 45.If the lateral wall angle cannot be found with confidence, the angle ofthe scanlines on the right side of the sector is used. The angle betweenthe broken lines 43,45 is bisected by a line 48, and the apex isinitially assumed to be located at some point on this line. With thehorizontal coordinate of the apex defined by line 48, a search is madeof the slope of pixel intensity changes along the line 48 to determinethe vertical coordinate of the apex. This search is made over a portionof line 48 which is at least a minimum depth and not greater than amaximum depth from the transducer probe, approximately the upperone-quarter of the length of line 48 above the mitral valve planebetween the MMA 26 and the LMA 36. Lines of pixels along the line 48 andparallel thereto are examined to find the maximum positive brightnessgradient from the LV chamber (where there are substantially no specularreflectors) to the heart wall (where many reflectors are located). Apreferred technique for finding this gradient is illustrated in FIG. 7.FIG. 7 a shows a portion of an ultrasound image including a section ofthe heart wall 50 represented by the brighter pixels in the image. Drawnnormal to the heart wall 50 is a line 48 which, from right to left,extends from the chamber of the LV into and through the heart wall 50.If the pixel values along line 48 are plotted graphically, they wouldappear as shown by curve 52 in FIG. 7 b, in which brighter pixels havegreater pixel values. The location of the endocardium is not the peak ofthe curve 52, which is in the vicinity of the center of the heart wall,but relates to the sense of the slope of the curve. The slope of thecurve 52 is therefore analyzed by computing the differential of thecurve 52 as shown by the curve 58 in FIG. 7 c. This differential curvehas a peak 56 which is the maximal negative slope at the outside of theheart wall (the epicardium). The peak 54, which is the first major peakencountered when proceeding from right to left along curve 58, is themaximal positive slope which is the approximate location of theendocardium. The pixels along and parallel to line 48 in FIG. 4 areanalyzed in this manner to find the endocardial wall and hence thelocation of the endocardial apex, marked by the small box 46 in FIG. 4.

If the user is operating on a sequence of stored images, the threepoints could be defined manually. For example, the user could point atthe three landmarks in an image of the sequence with a pointing devicesuch as a mouse or trackball, then click on them as they are identifiedto mark them in the image.

Once these three major landmarks of the LV have been located, one of anumber of predetermined standard shapes for the LV is fitted to thethree landmarks and the endocardial wall. Three such standard shapes areshown in FIGS. 5 a, 5 b, and 5 c. The first shape, border 62, is seen tobe relatively tall and curved to the left. The second shape, border 64,is seen to be relatively short and rounded. The third shape, border 66,is more triangular. Each of these standard shapes is scaledappropriately to fit the three landmarks 26,36,46. After anappropriately scaled standard shape is fit to the three landmarks, ananalysis is made of the degree to which the shape fits the border in theecho data. This may be done, for example, by measuring the distancesbetween the shape and the heart wall at points along the shape. Suchmeasurements are made along paths orthogonal to the shape and extendingfrom points along the shape. The heart wall may be detected using theoperation discussed in FIGS. 7 a-7 c, for instance. The shape which isassessed as having the closest fit to the border to be traced, by anaverage of the distance measurements, for instance, is chosen as theshape used in the continuation of the protocol.

The chosen shape is then fitted to the border to be traced by“stretching” the shape, in this example, to the endocardial wall. Thestretching is done by analyzing 48 lines of pixels evenly spaced aroundthe border and approximately normal to heart wall. The pixels along eachof the 48 lines are analyzed as shown in FIGS. 7 a-7 c to find theadjacent endocardial wall and the chosen shape is stretched to fit theendocardial wall. The baseline between points 26 and 36 is not fit tothe shape but is left as a straight line, as this is the nominal planeof the mitral valve. When the shape has been fit to points along theheart wall, the border tracing is smoothed and displayed over the endsystole image as shown in the image 78 on the right side of the dualdisplay of FIG. 8. The display includes five control points shown as X'salong the border between the MMA landmark and the apex, and five controlpoints also shown as X's along the border between the apex landmark andthe LMA landmark. In this example the portion of line 48 between theapex and the mitral valve plane is also shown, as adjusted by thestretching operation.

Since each of the images shown in FIG. 8 is one image in a cardiac loopof images, the clinician can further verify the accuracy of the bordersof the end diastole and end systole images 76,78 by playing a savedcardiac loop of images behind the borders drawn on the display of FIG.8. This is done by selecting one of the images of FIG. 8, then selecting“Play” from the system menu to repetitively play the saved cardiac loopin real time or at a selected frame rate of display behind the border.In the end diastole image 76 the endocardium is at its maximumexpansion; hence, the endocardium in the loop should appear to moveinward from and then back to the endocardial border drawn on the enddiastole image. In the end systole image 78 the endocardium is fullycontracted; hence, the endocardium in the loop should appear to moveoutward and then back to the border in this image. If the endocardiumdoes not move in this manner and, for example, is seen to pass throughthe border, a different image may need to be chosen for end diastole orend systole, or manual adjustment of a drawn border may be necessary. Ofcourse, the loop and its drawn borders over the complete cardiac cyclecan be replayed, enabling the clinician to view to endocardial tracingas it changes with the heart motion in real time.

Images with automatically traced borders which are saved in memory canbe recalled and their automatically drawn borders refined by manualadjustment, if desired. This process is knows as “rubberbanding.” AsFIG. 8 shows, the endocardial borders of both the end diastole and endsystole images have small boxes denoting the three major landmarks andcontrol points marked by X's on the septal and lateral borders. Theclinician chooses the default number of control point which will bedisplayed initially; on the border 80 shown in FIG. 9 there are threecontrol points shown on the septal wall and four control points shown onthe lateral wall. The clinician can review the end diastole and systoleimages, as well as all of the intervening images of the stored loop ifdesired, and manually adjust the positions of the landmark boxes andcontrol point X's if it is seen that the automated process placed aborder in an incorrect position. The clinician can slide a box or Xalong the border to a new position, and can add more control points ordelete control points from the border. Suppose that the ABD processorhad initially located the control point and border at the position shownby circle 82 and dashed line 84, which the clinician observes isincorrect. The clinician can relocate the control point laterally bydragging the X with a screen pointing device to the new location asshown by 86. As the X is dragged, the border moves or stretches alongwith the X, thereby defining a new border as shown by the solid lineborder 88. In this manner the clinician can manually correct and adjustthe borders drawn by the ABD processor.

As the ABD processor is identifying the key landmarks and fittingborders to the sequence of images, it is periodically making confidencemeasurements to gauge the likelihood that the image borders are beingaccurately located and traced. For instance, if the septum is notclearly contrasted from the blood pool in the LV chamber, the automatedprocess will stop. If the various correlation coefficients do not exceedpredetermined thresholds the process will stop. Both spatial andtemporal confidence measurements are employed. For instance, if thecomputed border of an image varies too much from a standard shape ineither size or shape, the process will abort. This can arise if thelandmarks are located in unusual positions in relation to each other,for example. If the change in the computed border from one image in thesequence to another is too great, the process will likewise abort. Whenthe process stops, a message is displayed notifying the clinician of thereason for stopping the process, and gives the clinician the option tocontinue the automated process, to continue the automated process withor after clinician input, or for the clinician to acquire a new loop ofimages or manually trace the current images.

In the illustrated example of FIG. 8 the automatically drawn borders ofthe end diastole and end systole images are used to compute the heart'sejection fraction. This is done by an automatic modified Simpson's ruleprocess which divides the delineated heart chamber at each phase into astack of virtual disks. The diameter of each disk is used with the diskheight to compute an effective volume of each disk, and these volumesare summed to compute the heart chamber volume at both end diastole andend systole. The difference between the two yields the ejectionfraction, the volume or percentage of the heart volume which is expelledas pumped blood during each heart cycle. The ejection fractioncalculation is shown in the measurement box at the lower left handcorner of FIG. 8 and is constantly updated. Thus, if the clinicianshould adjust a drawn border by the rubberbanding technique, thecomputed volume of the heart during that phase will change, affectingthe ejection fraction calculation, and the new calculation immediatelyappears in the measurement box. As the clinician adjusts the drawnborders he instantaneously sees the effects of these changes on thecalculation of the ejection fraction.

FIG. 10 illustrates biplane images which may be used in an embodiment ofthe present invention. As used herein the term “biplane images” refersto two images which are simultaneously acquired from different planes ofa volumetric region of the body. Two images are acquired simultaneouslywhen they are acquired in the same short time interval, which may beaccomplished by acquiring the first and second images in rapidsuccession or by interleaved acquisition of scanlines from the two imageplanes until the two images have been fully acquired. US patentapplication Ser. No. 10/231,704; allowed, Jul. 28, 2003 entitled“BIPLANE ULTRASONIC IMAGING”, incorporated herein by reference and ofwhich I am a co-inventor describes two biplane modes. In the biplaneimplementation described in this patent, one image is acquired in aplane normal to the center of a two dimensional array transducer and theother image is initially in a plane centered with the first image but ina plane orthogonal to that of the first image. One of the images maythen be relocated by either the “tilt” mode of operation or the “rotate”mode of operation. In the tilt mode, the second image is inclined toanother image plane which intersects another scanline of the first imagewhile remaining in an orthogonal orientation with the first image. Thesecond image may be tilted from alignment with the center of the firstimage to alignment with the lateral scanlines of the first image or inalignment with any scanline between these extremes. In the rotate modethe two images retain their common centerline alignment but the secondimage plane is rotated about this centerline. The second image can berotated from its initial orthogonal relationship with the first image tothe same image plane as the first image, or at any rotationtherebetween. FIG. 10 shows two biplane images in the rotate mode. Theleft image is a four-chamber heart image like those shown previously andthe right image is orthogonal to the first and shows the left ventricleas it appears when intersected by the plane of the second image. Thecircular white icon between the two images in the center of the screenshows that the right image plane has been rotated ninety degrees fromalignment with the left reference image plane. Marker dots are clearlyvisible in the icon and on the right sides of the apexes of the twosector images, indicating the left-right orientation of the two images.For completeness of a cardiac study the EKG trace is also shown belowthe biplane images.

An advantage of the present invention is that since only two planes of avolumetric region are being imaged, acquisition of the two images can bedone rapidly enough so that the two images can both be real-timeultrasonic images at a relatively high frame rate of display. Moreover,the two images are of the heart at substantially the same point in timeand are thus concurrently acquired images for purposes of the invention.A further advantage is that the ultrasound system need be only aconventional two dimensional imaging system. As FIG. 11 will illustrate,the display subsystem for biplane imaging can be a conventional twodimensional image processing subsystem, which means that biplane imagingin accordance with the present invention can be done with the twodimensional ultrasound systems currently in the hands of clinicians. Thescanner and display subsystem of FIG. 11 needs no unique 3D capabilitiesin order to produce the biplane image shown in FIG. 10.

Referring now to FIG. 11, an ultrasound system constructed in accordancewith the principles of the present invention is shown in block diagramform. In this embodiment an ultrasound probe 110 includes a twodimensional array transducer 500 and a micro-beamformer 502. Themicro-beamformer contains circuitry which controls the signals appliedto groups of elements (“patches”) of the array transducer 500 and doessome processing of the echo signals received by elements of each group.Micro-beamforming in the probe advantageously reduces the number ofconductors in the cable 503 between the probe and the ultrasound systemand is described in U.S. Pat. No. 5,997,479 (Savord et al.) and in U.S.Pat. No. 6,436,048 (Pesque).

The probe is coupled to the scanner 310 of the ultrasound system. Thescanner includes a beamform controller 312 which is responsive to a usercontrol and provides control signals to the microbeamformer 502instructing the probe as to the timing, frequency, direction andfocusing of transmit beams. The beamform controller also control thebeamforming of received echo signals by its coupling to theanalog-to-digital (A/D) converters 316 and the beamformer 116. Echosignals received by the probe are amplified by preamplifier and TGC(time gain control) circuitry 314 in the scanner, then digitized by theA/D converters 316. The digitized echo signals are then formed intobeams by a beamformer 116. The echo signals are then processed by animage processor 318 which performs digital filtering, B mode detection,and/or Doppler processing, and can also perform other signal processingsuch as harmonic separation, speckle reduction through frequencycompounding, and other desired image processing.

The echo signals produced by the scanner 310 are coupled to a displaysubsystem 320, which processes the echo signals for display in thedesired image format. The echo signals are processed by an image lineprocessor 322, which is capable of sampling the echo signals, splicingsegments of beams into complete line signals, and averaging line signalsfor signal-to-noise improvement or flow persistence. The image lines ofeach biplane image are scan converted into the desired image format by ascan converter 324 which performs R-theta conversion as is known in theart. The images are then stored side-by-side (see FIG. 10) in an imagememory 328 from which they can be displayed as one display frame on thedisplay 150. The images in memory are also overlayed with graphics to bedisplayed with the images, which are generated by a graphics generator330 which is responsive to a user control. Individual image frames orimage frame sequences can be stored in a cine memory 326 during captureof image loops.

For real-time volumetric imaging the display subsystem 320 also includesthe 3D image rendering processor 162 which receives image lines from theimage line processor 322 for the rendering of a real-time threedimensional image which is displayed on the display 150.

In accordance with the principles of the present invention, the biplanesystem includes an automatic border detection (ABD) processor 470 whichoperates as described above in real time to automatically trace themyocardial borders of the biplane images as they are produced. Theresult of border tracing of orthogonal biplane LV images is shown in theuser display of FIG. 12 which is that of a constructed embodiment of thepresent invention. In the embodiment of FIG. 12, the display screen isdivided into four quadrants. In the upper left quadrant one of thebiplane images is shown with the heart walls (endocardium 210 andepicardium 214) delineated by automatically drawn traces produced by theABD processor and overlaid over the ultrasound image by the graphicsgenerator 330. The orientation of the image plane of the ultrasoundimage in quadrant 202 is seen to be zero degrees.

A biplane image in a 90° orthogonal plane is shown in the upper rightquadrant 204. Like the first biplane image, the epicardium 216 andendocardium 212 of the LV have been delineated by automatically drawnborders.

A graphical model of the LV chamber volume produced as an interpolatedsurface spline 220 is shown in the lower left quadrant 206 of thedisplay. This surface spline is formed in this embodiment by fitting asurface to the orthogonal borders 210 and 212 as discussed below. Thesurface spline 220 encloses a volume which is measured by Simpson'sformula (rule of disks) to estimate the instantaneous capacity of the LVat each time of biplane image acquisition. These quantitative volumemeasures are displayed as a function of time in the lower right quadrant208 as illustrated by the physiologic curve 218 of the LV volume.Numeric measures of the volume at end diastole and end systole are shownto the right of the physiologic curve 218.

While various processes may be used to produce the spline surface 220 inFIG. 12, the technique used in a constructed embodiment is illustratedby FIGS. 13 a and 13 b. In the perspective view of FIG. 13 a the twotracings 210 and 212 of the endocardial border of the simultaneousbiplane images are shown on a base 220 which represents the mitral valveplane. The apex marker is shown at 230. In this example the image planesof the two biplane images are orthogonal to each other. The volumewithin the two tracings 210 and 212 is mathematically divided intospaced planes 222 which are parallel to the base plane 220. These planesintersect the left side of tracing 210 as shown at a,a and intersect theright side of tracing 210 as shown at c,c. The planes intersect the nearside of tracing 212 as shown at b,b.

An ellipse is mathematically fit to the four intersection points a,b,c,dof each plane 222 as shown in FIG. 13 b. While curves or splines otherthan ellipses can be used, including arcs and irregular shapes, anellipse provides the advantage that Simpson's formula has beenclinically validated when practiced with ellipses. FIG. 13 b shows anellipse 232 intersecting points a,b,c,d of the two tracings 210,212 nearthe base 220; an ellipse 234 intersecting the two tracings 210,212toward the center of the LV volume; and an ellipse 236 intersecting thetwo tracings 210,212 near the top of the LV volume. When ellipses havebeen fit on each of the spaced planes 222, the disk-like volumes betweenthe ellipses can be quantified by using the geometries of the adjacentellipses and the distances between them. The total volume of all of thedisk-like volumes is computed by summation to produce an instantaneousmeasure of the LV volume. The measure is used to produce another pointon the physiologic curve 218 and, if the ECG waveform indicates that theheart is at end systole or end diastole, the appropriate numerical valueis updated in the lower right quadrant 208 of the display.

In the constructed embodiment a surface is fit to the wire frame modelof ellipses and tracings. This can be done by interpolating a surfacethat fits smoothly over the wire frame. The graphical model 220 of thecavity volume is then displayed with a continuous surface.

In operation, the display of the embodiment of FIG. 12 appears asfollows. With the two dimensional array probe positioned to acquirebiplane images intersecting in the vicinity of the apex of the LV, realtime biplane images are displayed in the upper quadrants of the display.As each biplane image is acquired a corresponding border is traced oneach image by the ABD processor 470 and displayed with the image. Eachpair of time-concurrent tracings is used to produce an updated graphicalvolumetric model corresponding to the tracings of the biplane images atthat point in time. An updated quantified volumetric measure of the LVis displayed graphically and/or numerically in the lower right quadrantof the display. As the LV borders of the ultrasound images and theirtracings in quadrants 202 and 204 move inward and outward in real time,the graphical volumetric model 220 also swells and recedescorrespondingly in real time, and the physiologic curve in the lowerright quadrant scrolls in time as it is constantly updated.

It will be appreciated that other variations of the embodimentsdescribed above will readily occur to one skilled in the art. Forinstance, a wide variety of automatic or semi-automatic border detectionprocesses may be used, including those described in U.S. Pat. No.6,491,636 (Chenal et al.); U.S. Pat. No. 6,346,124 (Geiser et al.); andU.S. Pat. No. 6,106,465 (Napolitano et al.) While the biplane images inthe example of FIG. 12 are shown in orthogonal planes, other planarorientations than 90° may be used. This may be done, for instance, afterrotating one of the images as described in my aforementioned U.S. patentapplication Ser. No. 10/231,704; allowed, Jul. 28, 2003. Instead ofusing the images from two plane orientations, images from three or moreintersecting planes can be acquired, traced, and used to display orquantify volumes or volume measures. Acquiring more images for eachvolume determination will be more time consuming but can produce a moreanatomically accurate model of the cavity and more accurate volumemeasurements. Volumes can be defined globally or regionally. Forinstance, subregions on opposite sides of the LV can be defined byhemi-ellipses or splines and their volumes measured. Volumes whichchange over time can be identified and measured such as those producedby color kinesis. The volume of interest may be scanned using either a2D matrix array transducer or a mechanically rotating 1D arraytransducer. A variety of volume measures may be made, such as ejectionfraction and cardiac output determinations. The inventive technique canbe used to measure other volumetric objects in the body besides theheart, such as a suspected tumor or anatomy of a developing fetus.

1. A method for ultrasonically measuring a volume of a heart in realtime comprising: repetitively acquiring ultrasonic images of the heartduring a heart cycle in two intersecting image planes which extendthrough the heart in different directions at substantially the same timewith an ultrasound probe; using an automated processor to definecorresponding object borders in the ultrasonic images during the heartcycle; producing a plurality of quantified measures of the volume of theheart during the heart cycle from the defined object borders in thedifferent directions; and displaying measures of a continuous change inthe heart volume as the heart beats.
 2. The method of claim 1, furthercomprising producing a graphical model of the volumetric object usingthe defined object borders; and wherein producing quantified measuresfurther comprises producing quantified measures using the graphicalmodel.
 3. The method of claim 1, wherein displaying further comprisesproducing a display comprising real time images from the twointersecting image planes with a visually highlighted defined objectborder in each image and a quantified measure using the defined objectborder of the images.
 4. The method of claim 3, wherein producing adisplay comprising a quantified measure further comprises producing adisplay of changes in the volumetric object as a function of time. 5.The method of claim 3, wherein the display of changes in the volumetricobject as a function of time comprises a graphical display, a numericaldisplay or both a graphical and numeric display.
 6. The method of claim1, wherein acquiring ultrasonic images comprises acquiring ultrasonicimages of a chamber of the heart, wherein the corresponding objectborders comprise the wall of the chamber of the heart.
 7. The method ofclaim 2, further comprising producing a display comprising real timeimages from the two intersecting image planes with a visuallyhighlighted defined object border in each image, a real time graphicalmodel using the defined object borders, and a quantified measure usingthe defined object border of the images.
 8. The method of claim 2,wherein producing quantified measures further comprises using thegraphical model to produce a volumetric measure by the rule of disks. 9.The method of claim 2, wherein producing a graphical model comprisesfitting a series of curves to a wire frame structure formed by thedefined object borders.
 10. The method of claim 9, wherein the curvescomprise ellipses or hemi-ellipses.
 11. A method for ultrasonicallymeasuring the a volume of a heart comprising: acquiring a sequence ofultrasonic images of the heart in real time during a heart cycle in twointersecting image planes at substantially the same time with anultrasound probe, the intersecting image planes extending in differentdirections through the heart volume; using an automated processor todefine corresponding object borders in the ultrasonic images during theheart cycle; producing a real time graphical model of a volumetricregion of the heart using the defined object borders; and producing fromthe defined object borders a real time measure of a change in heartvolume during the heart cycle.
 12. The method of claim 11, wherein usingan automated processor further comprises using an automated processor toautomatically trace corresponding object borders in the ultrasonicimages; and wherein producing a graphical model comprises producing awireframe model by fitting a series of curves to the traces in theircorresponding image planes.
 13. The method of claim 12, wherein theseries of curves further comprise a series of ellipses.
 14. The methodof claim 12, wherein producing a graphical model further comprisesfitting a surface to the wireframe model.
 15. The method of claim 12,wherein producing a real time measure further comprises producingquantified measures of the graphical model by the rule of disks.
 16. Themethod of claim 11, further comprising producing a display comprisingreal time images from the two intersecting image planes with a visuallyhighlighted defined object border in each image and a real timegraphical model using the defined object borders.
 17. The method ofclaim 11, wherein acquiring comprises acquiring ultrasonic images of thevolumetric object in two or more intersecting image planes atsubstantially the same time with an ultrasound probe.